Fabrication of heart tubes from iPSC derived cardiomyocytes and human fibrinogen by rotating mold technology

Due to its structural and functional complexity the heart imposes immense physical, physiological and electromechanical challenges on the engineering of a biological replacement. Therefore, to come closer to clinical translation, the development of a simpler biological assist device is requested. Here, we demonstrate the fabrication of tubular cardiac constructs with substantial dimensions of 6 cm in length and 11 mm in diameter by combining human induced pluripotent stem cell-derived cardiomyocytes (iPSC-CMs) and human foreskin fibroblast (hFFs) in human fibrin employing a rotating mold technology. By centrifugal forces employed in the process a cell-dense layer was generated enabling a timely functional coupling of iPSC-CMs demonstrated by a transgenic calcium sensor, rhythmic tissue contractions, and responsiveness to electrical pacing. Adjusting the degree of remodeling as a function of hFF-content and inhibition of fibrinolysis resulted in stable tissue integrity for up to 5 weeks. The rotating mold device developed in frame of this work enabled the production of tubes with clinically relevant dimensions of up to 10 cm in length and 22 mm in diameter which—in combination with advanced bioreactor technology for controlled production of functional iPSC-derivatives—paves the way towards the clinical translation of a biological cardiac assist device.


A relatively high iPSC-CM density is required for achieving synchronized calcium transients in fibrin-based cardiac tubes
For initial experiments with cells the mold was shortened to 3 cm in length to enable efficient use of iPSC-CMs.The MHHi001-A-5 hiPSC reporter line ("Ruby"; constitutively expressing nuclear RedStar and the genetic calcium sensor GCaMP6f 25 ) was applied for the generation of CMs based on our chemically defined differentiation strategy in large-scale suspension culture 26,27 .This cell line enables both, evaluation of final cell density and monitoring of electrical coupling of hiPSC-CMs in living tissues.As the cell density is a crucial parameter for tissue formation and function, different concentrations of CMs were tested ranging from 5 × 10 6 -25 × 10 6 cells/ mL each composed of ~ 90% hiPSC-CMs and ~ 10% human foreskin fibroblasts (hFFs) (Fig. 3a-d).Notably, the addition of cells to the fibrinogen and thrombin solution even at the highest cell concentration tested did not impair the polymerization process and was fully compatible with proper tube formation (Fig. 3e).Furthermore, the centrifugal force of approximately 100 × g employed in the process did not hamper the cell viability as indicated by persistent nuclear RedStar fluorescence and active calcium signaling.The initial cell concentration applied correlates with the observation of RedStar positive cells quantified from images of the outer wall of the respective tubes (Fig. 3f).For the highest cell number tested, representing about 22.5 × 10 6 CMs/mL, a quantification of the cell density was not feasible as the multi-layered structure hindered an accurate discrimination of single cells.The highest cell concentration tested was notably required to retrieve tubes showing functional CMs coupling, as revealed by the synchronized oscillation of the GCamP6f signal 7 days after fabrication by RMT (Fig. 3g, h, supplemental movie1).

Rotating mold technology (RMT) promotes functional CM coupling
To evaluate the impact of RMT on cell distribution and functionality of cardiac tubes, our technology was compared to a "static casting" (SC) approach.Therefore, 3D printed casting molds were designed for generating tubes with an outer diameter of 8 mm and a wall thickness of 2 mm (Fig. 4a), closely reflecting dimensional properties of tubes generated by RMT (Fig. 4b).Applying the highest concentration of 25 × 10 6 cells/mL tested above for both technologies, revealed substantial differences regarding the cell distribution and functionality in the resulting tissues (Fig. 4).When observing the outer surface of SC-tubes, CMs were rather loosely distributed www.nature.com/scientificreports/accompanied by uncoordinated flashing activity of the calcium sensor in the isolated, individual CMs (Fig. 4c, e, g, supplemental movie 2).This distribution was confirmed over the entire wall of the fibrin tube in cross-sections (Fig. 4i).In contrast, when analyzing the outer surface of RMT-tubes after an equivalent time of tissue formation (i.e. 8 days after fabrication) a highly compact CM distribution was observed similar to the previous experiment (Fig. 3) showing a synchronized activity of GCamPf6 (Fig. 4d, f, h, supplemental movie 3).However, analysis of cross-sections revealed a sandwich-like structure of individual cell-dense layers, separated by essentially cell-free matrix (Fig. 4j).Quantifying the nuclei per area in these cross-sections from SC-tubes and RMT-tubes, here restricted to the cell-dense layer, confirmed the increase in cell density by the RMT process (Fig. 4k).The layered tissue characteristics in RMT-tubes may reflect the RMT dependent process properties, as RMT is based on moving the applicator dispensing the liquid cell-fibrinogen and cell-thrombin mixture, respectively, along the axis of the tube in repeated cycles.Nevertheless, our histological and functional analysis strongly indicates that RMT-induced formation of compact cell layers promotes the formation of physiologically coupled cardiac tissues.

Compaction of fibrin tubes depends on hFF content and medium composition
In previous research employing collagen and Matrigel™ as a matrix for the generation of sheet-like bioartificial cardiac tissues (BCTs), the constructs were cultivated in 12% horse serum (HS) containing "BCT medium" (BCTM) 17 .Employing such medium for fibrin tube cultivation resulted in a strong alteration/compaction of the 3D structure within 8 days, although the fibrinolytic inhibitor aprotinin (200 KIU/mL) was added.Such compaction, mainly as a result of fibrinolysis caused by hFFs, resulted in progressive tube disintegration while the alternative cultivation in a serum-free media formulation (SFM) 28 did not induce this process (data not shown).
To systematically investigate the impact of media composition, SC-tubes containing 22.5 × 10 6 Ruby-CMs/mL and 2.5 × 10 6 hFFs/mL were cut into rings of equal length and subjected to cultivation in BCTM and SFM with increasing amounts of HS (Fig. 5).Compaction-associated effects were observed for up to 7 days, revealing increased tissue shrinkage in reply to higher HS content in both media (Fig. 5a-c, d-f).Without HS addition, no shrinkage was observed irrespective of the basal medium (Fig. 5a, d).However, when combined with HS, the compaction was more pronounced in BCTM compared to SFM.Immunofluorescence staining of cross-sections revealed an increase in cell density in response to the compaction process (Fig. 5g-i).Quantification of nuclei per area in the respective cross-sections confirmed the increase of nuclei density upon compaction (Fig. 5j).This tissue compaction effect was notably accompanied by the synchronization of the Ca-sensor activity (data  not shown), corroborating the importance of proximal cell-cell contacts for the electrophysiological coupling of CMs in cardiac tissues.
The presence of fibroblasts is also known to play an important role in functional cardiac tissue formation in vitro 17 .Thus, the contribution of hFFs in the compaction process as well as fibrinolytic inhibition by t-AMCA or aprotinin supplementation was investigated by employing rings cut from SC-tubes (Fig. 6).The base medium contained 6% HS and 200 KIU/mL aprotinin.When entirely omitting hFFs from the cell-matrix composition, only minor compaction occurred, indicated by the reduction of cross-sectional area over time, independent of the inhibitors used (Fig. 6a), while hFF addition resulted in prominent, dose-dependent shrinkage (Fig. 6b, c).The addition of t-AMCA or doubling the concentration of aprotinin led to a comparable and slower degradation of the ring structures, resulting in a twofold larger cross-sectional area compared to BCTM controls on day 15.As a result of this systematic time-dependent investigation, it was decided to utilize a content of 10% hFFs for tube formation, followed by cultivation in BCTM containing t-AMCA in further experiments.

Conceptual re-design and scale-up results in a highly versatile device prototype for the controlled production of larger fibrin tubes
To overcome both technical limitations regarding the control of numerous process parameters related to RMT and limitations regarding the production of tubes with large diameters, a new machine was required.The new device prototype including the respective control software was designed and built together with the MHH research workshop (Fig. 7a-d).The set-up comprised the use of polyetheretherketone (PEEK)-based molds of variable dimensions, allowing the rotation mold production of tubes with 10 cm in length and a diameter of up to 22 mm, which was readily achieved in proof-of-concept experiments (Fig. 7e-g).However, by adjustment of mold dimensions, different lengths and diameters of tubes can be realized (Fig. 7h-j), whereby 6 cm in length and 11 mm in outer diameter were chosen for further experiments employing iPSC-CMs and hFFs.In addition, slight changes in the production process were implemented which were enabled by the new device.As noticed in previous experiments, cell loss was unavoidable due to centrifugal forces pushing cell-containing liquid out of the mold until sufficient polymerization of fibrin acted as a sealant.For an improved fabrication, a cell-free sheath was produced in the first step, directly followed by the application of iPSC-CMs and hFFs containing fibrinogen and thrombin solution forming a tube-in-tube structure with the possibility to remove the outer sheath (Fig. 7j).www.nature.com/scientificreports/

Tube Production with the new device combined with the sheath strategy leads to functional cardiac tubes of substantial dimensions
To ensure the general applicability of our method, an additional iPSC line MHHi001-A 29 was used.In addition, an analysis of bioartificial cardiac tissue (BCT) generated by a different technique 30 revealed that force generation of these tissues depends on the iPSC cell lines used for CMs production.Tissues derived from CMs originating from the initially employed iPSC cell line MHHi001-A-5 exhibited lower forces compared to BCTs derived from the parental iPSC line MHHi001-A (unpublished data).As the iPSC line MHHi001-A does not carry any reporter, cell viability was demonstrated by TMRM staining and further applied throughout the entire cultivation time for imaging of vital cells.RMT-tubes (6 cm in length, 11 mm outer diameter) produced with an outer sheath were equipped with connectors and subjected to pulsatile flow of 10 mL/min in a bioreactor at the frequency of 1 Hz (Fig. 8a, b, supplemental movie 4).Employing 10% of hFFs with a total number of 150 × 10 6 cells per tube and cultivation in BCTM with t-AMCA enabled prolonged cultivation of up to 5 weeks without substantial shrinkage of the fibrin tube.TMRM staining revealed the even distribution of cells in circumferential and longitudinal direction (Fig. 8c-e).After discontinuing the pulsatile flow, spontaneous contractions of the fibrin tubes were visible.Contractions were recorded at different positions and analyzed revealing 65 ± 9 beats per minute (Fig. 8f).
Further functionality of the iPSC-CMs was demonstrated by their proper reaction to electrical stimulation via pacing at a frequency of 2 Hz (Fig. 8g).Staining of cross-sections revealed an outer cell-dense layer of CMs and fibroblasts (Fig. 8h) with CMs exhibiting cross-striations (Fig. 8i).Deposition of collagen I was demonstrated revealing intense staining for collagen when co-localized with vimentin expressing fibroblasts (Fig. 8j, k) while showing faint staining in regions negative for vimentin.A cell density of 51 ± 7.5 nuclei/10,000 µm 2 was determined by automatic counting of cross-sections; this value is slightly higher compared to the cell-dense layers in the RMT-tubes produced in the previous set-up (Fig. 4k).

Discussion
Here, we have demonstrated the feasibility to produce large, tubular fibrin constructs based on commercially available human fibrinogen and thrombin employing iPSC-CMs using RMT.Although the use of fibrinogen from FFP seems reasonable, inhomogeneous fibrin tubes were produced from cryoprecipitated fibrinogen.The isolation by cryoprecipitaion results presumably in fibrinogen with a high content of impurities, which might alter the rheological behavior leading to an uneven distribution along the axis of the mold and therefore resulting in the observed heterogeneity in Young's modulus along individual tubes (Fig. 2).Hence, commercially available fibrinogen (cFb) significantly increased the experimental reproducibility resulting in rather homogenous tube formation by RMT, cFb was employed in further experiments.Through the centrifugal forces in the fabrication process CMs are concentrated and brought into close proximity to each other facilitating the coupling of CMs as early as day 3 of cultivation as indicated by the activity of the calcium sensor GCaMP6f (Fig. 3, supplemental movie 1).Comparison of RMT-tubes to SC-tubes employing the same cell concentration of 25 × 10 6 cells/mL clearly demonstrated the increase of cell density in the rotation process (Fig. 4).The final cell density is an important aspect for functional tissue formation as local cell-cell communication via a variety of secreted factors as well as direct CM-CM communication via gap junctions is crucial for the maturation of CMs and impulse conduction 31 .In the final set-up the applied cell concentration was even increased to 50 × 10 6 cells/mL reaching the technical limit of the set-up as higher cell densities impaired the transport of solutions via the drip line (data not shown).Further analysis of the impact of the ratio of hFFs as well as medium components on remodeling of the fibrin matrix revealed that a combination of 10% hFFs and the addition of t-AMCA results in moderate compaction enabling the culture of fibrin tubes for up to 5 weeks without losing their initial structure.Aiming at the future implementation of iPSC-derived cardiac fibroblasts, this data would need to be revised as Boucard et al. reported the dependence of fibrin degradation on fibroblast origin 32 .They observed that proteases such as MMP1, MMP3, www.nature.com/scientificreports/and urokinase, which are directly or indirectly involved in fibrinolysis, were expressed at varying levels depending on the derivation of fibroblasts 32 .In the same study, it was demonstrated that aprotinin exhibited a dosedependent effect on degradation, which was reduced by a factor of 2 when increasing the aprotinin concentration from 20 µg/mL to 40 µg/ml 32 .A similar trend was observed in our study when increasing the aprotinin content from 200 KIU/ml to 400 KIU/mL in the ring-based assay (Fig. 6).Prevention of fibrin degradation is not only important in vitro but of utmost importance in vivo as native fibrin will be degraded as soon as it is implanted into the body.Engineering a fibrin-based construct suitable for transplantation requires either the modification of fibrin per se, e.g.via crosslinking, or the replacement of fibrin by extracellular matrix proteins deposited by cells during the in vitro cultivation.Indeed, Syedain et al. generated implantable tissue-engineered blood vessels as well as heart valves by prolonged cultivation of fibroblasts replacing the initial fibrin matrix by collagen 33,34 .In our study, collagen deposition by hFFs and also by CMs was demonstrated (Fig. 8) but further analysis would be required to determine if the retrieved tissue exhibits sufficient stability.Collagen deposition could be increased by activation of fibroblasts or by increasing their amount with the drawback of increased fibrinolytic activity as demonstrated in our study (Fig. 6).These findings also highlight the importance of thoroughly fine tuning the fibrinolytic and anti-fibrinolytic components in fibrin-based cardiac tissue engineering.
With future clinical applications in mind, we have employed a rather high fibrinogen concentration of 30 mg/ mL aiming at establishing a mechanically relatively strong matrix.Nanoindentation revealed a Young's modulus of 2.6 ± 0.5 kPa on average, which is well in line with a value of 1.0 ± 0.3 kPa for fibrin based hydrogels with a concentration of 15 mg/mL fibrinogen reported by Jung et al. employing a similar nanoindenter 35 .Jung et al. observed a large scattering of the values retrieved by nanoindentation, which we detected in our data sets as well.The high deviation of the values seems to be inherent to the technology i.e. measuring fibrous samples with probes in the micrometer range.Although the Young's modulus still resembles values for rather soft tissues, tubes could be secured with a surgical thread onto plastic connectors for implementation in a perfusion bioreactor and could withstand subsequent stimulation by pulsatile flow (Fig. 8).
In a recent publication Koehne et al. followed a casting strategy for the generation of tubular heart tissue employing bovine fibrinogen with a final concentration of 10 mg/mL.Koehne et al. employed a silicon tubing as well as velcro rings to secure tube integrity 12 .In addition, the necessity for a silicon support tube hindered the assessment of burst pressure in their setting.Also others used a tubular support such as an octagonal column 9 or a decellularized human umbilical artery 10 for cardiac tubular tissue engineering.In these approaches CMs were introduced as cell sheets generated by culture on thermoresponsive tissue culture plates 9 or cultured on decellularized slices of porcine ventricular heart tissue 10 wrapped around the tubular support.Williams et al. used thermoresponsive nanofabricated substrates with cell-sheet stacks wrapped into a cylindric shape which is stabilized by casting and crosslinking a hydrogel around the sheets 11 .All these studies have in common that a technique was implemented to produce a cell dense layer to foster the coupling and functionality of employed CMs.In the current study, this is easily achieved by applying a high density of up to 50 × 10 6 cells/mL within the rotation process without labour intensive pre-cultivation of cell sheets.Remarkably, the initial cell density used in our study is much higher compared to reports by Koehne et al. 12 (18 × 10 6 cells/mL) and Querdel et al. 36 (15 × 10 6 cells/mL) both following a casting approach.Through the application of centrifugal forces inherent in our approach, the cells are in addition even more concentrated at the outer perimeter of the tube probably approaching the cell density of the native human heart with 100 × 10 6 cells/mL 37 .Functionality of our fibrinbased cardiac tubes was demonstrated via their ability to spontaneously contract with 65 ± 9 beats per minute correlating well with the mean real-world heart rate of 79 ± 15 beats per minute recently determined in the eHeart study 38 .Furthermore, the tissues were able to react to electrical stimuli with increasing their beating frequency according to the pacing (Fig. 8).
The reported survival and functionality of iPSC derived CMs has been demonstrated for up to 100 days for hiPSC derived 3D cardiac aggregates 39 and organoids 40 in vitro.In vivo, the presence of transplanted hiPSC-CM has been revealed for at least 3 months in large animal models shown by us 41 and others 42 , providing meaningful hints for the long-term engraftment and phenotypic stability of iPSC-CMs and thus supporting the envisioned therapeutic concept.So far, the described approaches have resulted in rather small constructs with 1.8 cm in length and 6 mm inner diameter published by Koehne et al. being the largest one reported 12 .The current dimensions of our approach may not yet fulfill the clinical requirements for a biological cardiac assist device, but our constructs are generally suitable for first planned functionality tests in ex vivo perfused porcine hearts.As our approach is straightforward regarding the further scalability to larger formats (Fig. 7), dimensions suitable for preclinical testing in animal models can be achieved.Clearly, with the increase in size, nourishment of cellcontaining constructs represents a major challenge that needs to be addressed.However, in combination with the advanced upscalable bioreactor technology for the GMP-compliant production of large amounts of iPSC derivatives 26,43 the presented technology paves the way for the production of functional cardiac tubes for clinical translation envisioned as biological cardiac assist device.

Fibrinogen isolation
Fresh frozen plasma concentrates (FFP) unsuitable for clinical use were obtained from the blood bank at Hannover Medical School.Frozen plasma was thawed at room temperature and centrifuged at 800 × g for 12 min.The supernatant was discarded and the hydrous protein pellet was solubilized at 37 °C.www.nature.com/scientificreports/

Fibrin tube production
Initially, fibrin tubes were fabricated in a machine established by Aper et al. 23 with a few modifications.A schematic of the process is depicted in Fig. 1.In brief, a rotating mold was driven by an electric motor.The apparatus consisted of an outer brass tube with 10 pairs of drill holes (diameter 0.3 mm, distance 1 cm) allowing the drainage of excess fluid, and two removable polyetheretherketone (PEEK) half shells inside.The cylindric mold was closed on both sides with PEEK stoppers with the front stopper provided with a hole enabling the insertion of an applicator.The applicator was driven by an electric motor enabling the movement in the mold along the axis.Two cannulas connected via drip lines to syringes driven by injectomats were inserted into the applicator.Thrombin and fibrinogen solutions were applied through these cannulas at a defined flow rate into the mold while the applicator was moved along the axis.As this apparatus allowed only a limited diameter for the tube and had limited control over movement speed and other parameters, a more sophisticated machine was built enabling the fabrication of fibrin tubes up to 34 mm in diameter (Fig. 6).Fully automated control over application speed off the solutions, movement of the applicator, and rotation speed was implemented allowing the execution of complex fabrication protocols.
For the fabrication of fibrin tubes two solutions were prepared independently.Commercially available human fibrinogen (Merck) and commercially available human thrombin (Merck) were solubilized in 0.9% NaCl.CaCl 2 solution was added to the thrombin solution.Application of equal volumes of the two solutions resulted in final concentrations of 30 mg/mL fibrinogen, 15 U/mL thrombin, and 40 mM CaCl 2 .Process parameters for the old machine were the following: centrifugal force of 72-162 × g (strong variation due to the properties of the motor), application speed for both solutions of 1.25 mL/min, and applicator speed of 1 mm/s.A volume of 3 mL for each solution was applied.Process parameters for the new device were the following: centrifugal force of 125 × g, application speed for both solutions of 1 mL/min, and applicator speed of 1.33 mm/s.A total volume of 4.5 mL was applied for the fabrication of a cell-free sheath and a total volume of 3.5 mL was used for cell application.

Nanoindentation
Pieces from the fabricated fibrin tubes were glued to a glass bottom dish, incubated for 30 min in PBS, and subjected to nanoindentation with PIUMA nanoindenter (Optics11, Amsterdam, The Netherlands) employing a 0.29 N/m cantilever-based probe with a tip radius of 49 µm.Measurements were done at room temperature.Per sample, 100 indentations were performed in a grid of 1 × 1 mm with a distance of 100 µm between individual indentations.The displacement profile is set to displace 20 µm from the automated near-surface definition.The data in the loading section of the load-displacement curve was used to determine the Young's Modulus using a fit of all data points from the contact point to 7 µm indentation depth respecting the Hertz model assumption of a parabolic indenter (max.depth for calculation equals to 16% of tip radius).Two to three pieces per tube were analyzed.

Figure 1 .
Figure 1.Schematic of the rotating mold technology.Two syringes containing fibrinogen solution and thrombin solution, respectively, are each connected to a drip line and a cannula.The cannulas are inserted into an applicator, which is reaching into a rotating mold and can move backward and forward.While moving, the two solutions are dispersed from the tip of the cannulas into the rotating mold.Upon contact of the two solutions the formation of fibrin fibers is initated.(Figure created with BioRender.com)

Figure 2 .
Figure 2. Commercially available fibrinogen improves uniformity of fibrin tubes.(a,b) Fibrin tubes fabricated with fibrinogen isolated from two independent batches of fresh frozen plasma (FFP), termed FFP1 and FFP2, respectively.(c) Tube produced from commercially available fibrinogen (cFb).(d) Young's modulus of tubes shown in (a-c).Young's modulus of pieces from position 1 and 2 (P1, P2) was determined by nanoindentation (3 pieces per tube were assessed by a matrix scan of 100 indentations per piece).Mean value with SD is displayed.(e) Young's modulus determined by nanoindentation of five independently fabricated tubes made from cFb. 2 pieces per tube were analyzed with > 100 measurement points per piece.Mean value with SD is displayed.ns = p > 0.05, ** = p < 0.01, *** = p < 0.001.FFP: fresh frozen plasma, cFb: commercially available fibrinogen.Scale bar: 1 cm.

Figure 3 .
Figure 3. High density of iPSC-CMs leads to synchronized calcium transients.(a-d) Fibrin tubes fabricated with increasing cell concentrations of Ruby-derived CMs and hFFs ((a) 5 × 10 6 /mL, (b) 10 × 10 6 /mL, (c) 15 × 10 6 / mL, (d) 25 × 10 6 /mL of total cell concentration with 90% CMs and 10% hFFs) after 3 days of cultivation in serum-free media (SFM).The indicated cell concentrations refer to the cell-matrix solution applied for the tissue generation process.(a-d) View onto the outer tube surface revealing nuclear RedStar fluorescence.(e) Mold for generating cell containing fibrin tubes.Length was reduced to 3 cm.Tube from (d) is shown below the mold directly after the fabrication.(f) Quantification of cells per region of interest (ROI) in dependence of initially applied cell concentration.Four ROIs with the same dimensions were analyzed per tube.Mean and SD is depicted.***= p < 0.001 (g) Still images of GCaMP6f signal recording of the tube shown in (d) displaying a time interval of 0.25 s between pictures I to VI demonstrating GCaMP6f activity by GFP signal intensity at day 7 of cultivation.(h) Calcium fluctuations represented by GFP signal intensity were analyzed over time with ImageJ employing the recording shown in supplemental movie 1.Four regions of interest (ROI) were randomly selected.Scale bar: (a-d,g) 500 µm, (e) 1 cm.

Figure 4 .
Figure 4. Rotation mold technology (RMT) facilitates CM coupling in contrast to static casting (SC) resulting in isolated CMs.(a,c,e,g,i) SC-tube.(b,d,f,h,j) RMT-tube.Tubes were fabricated employing the same cell concentration.(a) Macroscopic appearance of SC-tube.Left inset: 3D printed mold for static casting.Right inset: Lumen of SC-tube.(b) Macroscopic appearance of RMT-tube.Left inset: Machine for fabrication (copyright: Karin Kaiser/MHH).Right inset: Lumen of RMTtube.(c, d) View onto the outer tube surface revealing nuclear RedStar fluorescence after 8 days of cultivation in SFM.(e, f) GCaMP6f peak fluorescence of Ruby CMs after 8 days of cultivation in SFM.Still images from movie.(g, h) Calcium fluctuations represented by GFP signal intensity were analyzed over time with ImageJ employing the recording shown in supplemental movie 2 and 3, respectively.Four regions of interest (ROI) were randomly selected.(i, j) Immunofluorescence of cross-sections stained for sarcomeric α-actinin (αSA, green), vimentin (red), and DAPI (blue).The dashed line represents the outer surface of the fibrin tube.(k) Quantification of nuclei per area for cross-sections from SC-tubes and RMT-tubes, respectively.Four regions of interest (ROI) were selected.For SC-tubes ROI were selected randomly, for RMT-tubes areas in the cell dense layers were selected.Mean and SD is depicted.*** = p < 0.001.Scale bar: (a,b) 1 cm for inset and main picture, (c-f) 200 µm, (i,j) 100 µm.

Figure 7 .
Figure 7. Automated machine for the production of fibrin tubes.(a) View of the machine.(b) Stainless steel pipe fitted on the motor for rotation.(c) Applicator equipped with 2 cannulas connected to drip lines.The motor for moving the applicator is located in the back.(d) Automated syringe feed with valves for closing and opening the lines.(e) PEEK mold for tube production.(f) Fibrin tube with a length of 10 cm and a diameter of 22 mm produced with the machine.Top view with one half shell removed.(g) Side view into the lumen of the tube in (f).(h) Fibrin tube produced in a mold with a length of 6 cm and an outer diameter of 11 mm.Top view with one half shell removed.(i) View into the lumen of a tube in the mold.(j) (I) Ring cut off from a tube produced with an outer sheath.(II) Same Ring as in (I) after the removal of the sheath which was cut open and placed on the left side of the ring.Scale bar: (f-j) 1 cm.

Figure 8 .
Figure 8. Analysis of fibrin tubes fabricated with the new device and cultivated under pulsatile flow.(a) Bioreactor set-up for pulsatile flow.(b) Fibrin tube mounted into a bioreactor.(c) View into the lumen of the tube stained with TMRM.(d) Side view of a tube stained with TMRM.Four images were stitched together.(e) Picture with high magnification of the tube in (d).(f,g) Videos from contracting fibrin tubes stained with TMRM were recorded, imported into ImageJ and analyzed with the MYOCYTER macro to reveal the beating frequency.The amplitude is depicted in black, the maxima in red and the minima in green.(f) Analysis of spontaneous beating fibrin tube.(g) Analysis of tube from (f) paced at 2 Hz.(h-k) Immunofluorescence of cross-sections.(h,i) Staining for sarcomeric α-actinin (αSA, green), vimentin (red), and DAPI (blue).(j,k) Staining for collagen I (green), vimentin (red), and DAPI (blue).Scale bar: (c,d) 1 mm, (e) 500 µm, (h,j) 50 µm, (i,k) 20 µm. https://doi.org/10.1038/s41598-024-64022-7